Radiological imaging apparatus and method

ABSTRACT

A radiological imaging apparatus and method corrects fixed pattern noise (FPN) generated within the apparatus and suppresses increases in random noise attendant upon FPN correction to provide radiological imaging having improved reliability and image quality. The apparatus has a signal processing circuit that digitizes analog output from a read-out circuit, four FPN memories for storing four dark outputs, a light memory for storing a light output after X-ray exposure of a subject, a CPU for controlling the signal processing circuit as well as FPN and light memories, and a shift resister  7  controlled by the CPU. The method involves averaging multiple dark outputs and subtracting the FPN data average so obtained from the light output which includes the X-ray imaged to enhance picture quality. Since the FPN correction uses FPN data that has been averaged over multiple dark outputs, the method also suppresses random noise generated by the FPN correction for enhanced imaging accuracy.

FIELD OF THE INVENTION

The present invention relates to a radiological imaging apparatus andmethod, and more particularly, to a radiological imaging apparatus andmethod used in medical X-ray equipment, industrial non-destructivetesting devices and the like.

BACKGROUND OF THE INVENTION

With the conventional still-picture X-ray technology currently in use,the so-called film-based technique is the most prevalent. In thistechnique, the patient is exposed to X-rays and the X-rays that passthrough the body are then exposed onto a sheet of film. The film has thefunction of displaying and recording information, and is widely usedthroughout the world due to its capacity to be enlarged, its high degreeof gradation, its light weight and ease of handling. On the other hand,the technique suffers from several disadvantages, including acomplicated process of developing the image, the problem of long-termstorage, and the time and effort involved in manual search and retrievalof the physical images.

By contrast, moving-image photographic systems rely mainly on imageintensifier (I.I). Since I.I uses the photo-electron multiplier effectinside the device, it generally has good sensitivity and has theadditional advantage of exposing the patient to lower levels ofradiation. The I.I technique not only provides the physician with asee-through image of the patient but also, due to the conversion of theCCD analog output to digital output (a process referred to here asdigitization), makes possible the computerized recording, display andstorage of such data. However, because medical diagnosis requires a highdegree of gradation, even with I.I, film is often used for still pictureimaging.

Recently, with a growing need to digitize X-ray images inside thehospital itself, in place of film, X-ray imaging devices that use anX-ray sensor with solid-state image sensing elements arrayedtwo-dimensionally to convert the X-ray image into electrical signalshave begun to be used. Since the X-ray image can then be replaced withdigital information, image information can be sent instantaneously todistant locations, with the advantage of being able to providestate-of-the-art, high-quality diagnostics even to remote areas.Moreover, if no film is used the space previously required for itsstorage can be turned to other, more productive uses. If in the futureit becomes possible to introduce more advanced and sophisticated imageprocessing techniques, it is possible that diagnostics may to someextent be computerized and therefore automated, without the interventionof a radiologist.

Moreover, in recent years, with the use of amorphous thin-filmsemiconductors in solid-state image sensing elements, X-ray imagingdevices capable of taking still pictures have been developed. Usingamorphous silicon thin-film semiconductor production technology, photosexceeding 40 cm a side and capable of completely imaging the human torsohave been commercialized. Since the production process itself isrelatively simple, it is expected that inexpensive detectors based onthis technique will become available in the not-so-distant future. Inaddition, since amorphous silicon can be produced in thin glass sheetshaving a thickness of i mm or less, the detector itself can be made verythin and compact, for greater ease of handling.

An X-ray imaging device of the sort described above typically has aphotoelectric converter circuit, on which a plurality of photoelectricconverters for converting radiation into electrical signals are arrangedas in a matrix, and a read-out circuit for reading out the electricalsignals form the photoelectric converter circuit. FIG. 7, for example,is a timing chart showing the operation of an X-ray imaging devicehaving nine photoelectric converters S1-1 through S3-3 (includingswitches (TFT) T1-1-T3-3). As for the photoelectric conversion interval(given in the diagram as the X-ray exposure interval): In a state inwhich the TFT are all OFF, when the light source (X-rays) are turned ONin pulses, the respective photoelectric converters are struck by thelight and a signal electric charge comparable to the amount of light isstored in the converter capacitor. If a fluorescent material is used toconvert the X-rays into visible light, then either a light-guidingmember for guiding the light made visible in proportion to the number ofX-rays to the photoelectric converters may be used or the fluorescentmaterial may be disposed near the electrodes of the photoelectricconverters.

It should be noted that the signal electrical charge is held in theconverter capacitor after the light source is OFF.

Next, as for the read-out interval: The read-out is accomplished at theS1-1-S3-3, one row at a time, starting with row S1-1-S1-3, then with rowS2-1-S2-3, and finally with row S3-1-S3-3. First, a gate pulse isapplied from SR1 to the switches (TFT) T1-1-T1-3 gate lines in order toread out the first row S1-1-S1-3. Doing so turns T1-1-T1-3 ON and thesignal electrical charge that had been stored in S1-1-S1-3 is sent tothe signal lines M1-M3 to which read-out capacitors CM1-CM3 have beenadded, so that the signal electrical charge is sent to the capacitorsCM1-CM3 via the TFT. For example, read-out capacitor CM1 added to signalline M1 is the (three-)sum total of the T1-1-T1-3 gate-sourceinterelectrode capacitance (Cgs). Amps A1-A3 amplify the signalelectrical charge sent to signal lines M1-M3.

The amplified signal electrical charge sent to capacitors CL1-CL3 bothturns OFF and holds SMPL signal OFF. Next, by imparting a pulse from ashift resister 103 to switches Sr1, Sr2 and Sr3 (in that order) thesignals held at CL1-CL3 are then output from an amp 104 in the orderCL1, CL2 and CL3. Since analog signal outputs B1, B2 and B3 are outputfrom the amp 104, the entire unit, including the shift resister 103 andthe switches Sr1-Sr3, is called an analog multiplexer. Ultimately, onerow's worth of photoelectric conversion signals (S-1, S1-2, S1-3) isoutput in sequence by the analog multiplexer. The read-out of the secondrow S2-1-S2-3 is carried out in the same way as the read-out of thefirst row described above.

If the signals at signals lines M1-M3 are sampled and held at CL1-CL3 bythe first row's SMPL signal, then the signal lines M1-M3 can be reset toground electric potential by a CRES signal and thereafter a G2 gatepulse can be applied. In other words, second-row signal electricalcharges from the photoelectric converters S2-1-S2-3 can be transmittedby the SR1 while at the same time the first row's signals are beingserially converted by the SR2. In so doing, all the signal electricalcharges of the first through third rows of photoelectric converters canbe output.

The photoelectric converter circuit operation described above is capableof reading X-ray images.

But in reality the image as such also contains fixed pattern noisegenerated by the photoelectric converter circuit and the read-outcircuit.

There are a number of factors that cause fixed pattern noise, including(A) variations in photoelectric converter dark current, (B) variationsin switch leak current, (C) variations in photoelectric convertercircuit signal wire capacitors CM1-CM3 and (D) variations in read-outcircuit amp (for example A1-A3) offset voltage.

Whereas causes (A) and (B) described above show up as dots in the finalimage, causes (C) and (D) appear as lines; in each case, they degradethe quality of the image. Therefore, conventionally, this type of fixedpattern noise is read out as a dark output image, which is thensubtracted from the X-ray image so as to correct the fixed pattern noise(hereinafter FPN).

The dark output image itself acquired in one of two ways: it is eitherread in when the unit is shipped from the factory or prior tophotographing and then stored in memory as dark output data, or it isobtained just before or just after read-in of an X-ray image.

However, there are problems with the conventional methods of acquiringthe dark output image (FPN).

In the former case, when the FPN is set either at shipment or prior tophotography, there are two problems:

-   -   (a) The dark currents of the photoelectric converters have        temperature characteristics, which may be different when the FPN        is obtained and when the X-ray image is shot. Such differences        in temperature characteristics can prevent the FPN from being        fully corrected.    -   (b) Since the TFT leak current changes over time, the leak        current may be different when the FPN is obtained and when the        X-ray image is shot, and such differences in leak current can        prevent the FPN from being fully corrected.

The technique of reading in the FPN just before or just after taking theX-ray image, although it solves the two problems described above, has aproblem of its own:

-   -   (c) random noise generated by the photoelectric converter        circuit (hereinafter referred to simply as random noise) as well        as random noise generated by the read-out circuit, which is not        fixed pattern noise and therefore cannot be fully corrected, and        which, when processed for elimination, only increases by a        factor of {square root}2 and degrades the quality of the image.

Random noise is caused by shot noise from the photoelectric converterdark electrical charge and by heat noise (that is, Johnson noise) causedby electron thermal disturbances inside the switches. It is one type ofnoise that is unavoidable. In addition, if the signal lines M1-M3 haveany appreciable internal resistance, that resistance, too, will generateJohnson noise or random noise. The operational amplifier that forms partof the read-out circuit also generates random noise.

The amount or volume of random noise can be determined by focusing on aparticular pixel and sampling its dark data several times in order toobtain its dark data distribution. In other words, that distribution isa Gaussian distribution, i.e., noise that has a frequency distributionthat follows the Gaussian curve, and can be calculated by obtaining thestandard deviation a.

The dark image output from an X-ray imaging apparatus comprising thephotoelectric converter circuit in which the photoelectric converter arearranged in a two-dimensional array, contains both FPN as well as randomnoise. Moreover, both types of noise appear not only in the dark imageoutput but also in the X-ray image as well. Subtracting the dark imageoutput from the X-ray image cancels the FPN but leaves the random noiseuncorrected. Also, between the dark image output and the X-ray image,the random noise itself can differ from one pixel to the next. Therandom noise generated when reading in the dark image output and therandom noise generated when reading in the X-ray image are mutuallyindependent phenomena, so an image in which the former is subtractedfrom the latter (that is, FPN correction) will have a volume of randomnoise ({square root}2·σ) that is {square root}2 the volume of randomnoise before correction (standard deviation σ).

SUMMARY OF THE INVENTION

The present invention has been conceived in light of the foregoingconsiderations.

According to one aspect of the present invention, preferably, aradiological imaging apparatus for forming X-ray images comprises: aradiation detection unit composed of a plurality of photoelectricconverters arrayed in a matrix-like formation for the purpose ofdetecting and converting radiation from an irradiated subject intoelectrical signals; a drive control unit operatively connected to theradiation detection unit for the purpose of controlling movement of theradiation detection unit; a read-out unit operatively connected to theradiation detection unit for reading the electrical signals output bythe radiation detection unit; and a signal processing unit operativelyconnected to the read-out unit and the drive control unit for processingthe electrical signals read out by the read-out unit, the signalprocessing unit selecting between a first read-out operation, in whichthe radiation detection unit is operated with irradiation of a subject,and a second read-out operation, in which the radiation detection unitis operated without irradiation of the subject, the signal processingunit further performing the second read-out operation at least twice,averaging the results of the second read-out operation so as to obtainan average signal value thereof, and subtracting the average signalvalue from a signal value obtained by the first read-out operation so asto obtain a processed image.

According to another aspect of the present invention, preferably, aradiological imaging method comprises the steps of: detecting andconverting radiation from and irradiated subject into electrical signalsusing a photoelectric converter unit; reading the electrical signalsoutput by the photoelectric converter unit using a read-out unitoperatively connected to the photoelectric converter unit; processingthe electrical signals connected to the electrical signal read-out unitso as to select between a first read-out operation, in which the subjectis irradiated with radiation, and a second read-out operation, in whichthe subject is not irradiated; performing the second read-out operationat least twice and averaging the results of the second read-outoperation so as to obtain an average signal value thereof; andsubtracting the average signal value from a signal value obtained by thefirst read-out operation so as to obtain a processed image.

Other features, effects and advantages of the present invention will beapparent from the following description, taken in conjunction with theaccompanying drawings, in which like reference characters designate thesame or similar parts throughout the figures thereof.

BRIEF DESCRIPTION OF THE DRAWINGS

The accompanying drawings, which are incorporated in and constitute apart of the specification, illustrate embodiments of the invention and,together with the description, serve to explain the principles of theinvention, in which:

FIG. 1 is a circuit diagram showing schematically the structure of anX-ray imaging apparatus using an amorphous silicon thin filmsemiconductor according to one embodiment of the present invention;

FIG. 2 is a circuit diagram showing the internal structure of theread-out circuit of FIG. 1;

FIG. 3 is a schematic diagram showing the internal structure of thesignal processing circuit of the X-ray imaging apparatus of FIG. 1;

FIG. 4 is a timing chart of the operation of an imaging apparatusaccording to one embodiment of the present invention;

FIG. 5 is a schematic diagram showing steps in a process of averagingtwo continuous pieces of FPN data according to one embodiment of thepresent invention;

FIG. 6 is a schematic diagram showing steps in a process of averagingtwo continuous pieces of FPN data according to a variation of anembodiment of the present invention;

FIG. 7 is a timing chart showing the operation of a conventional X-rayimaging device; and

FIG. 8 shows actual measurements of noise level versus number ofaveraging processes using the X-ray imaging apparatus.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

Preferred embodiments of the present invention are described in detailin accordance with the accompanying drawings, in the first instance FIG.1.

FIG. 1 is a circuit diagram showing schematically the structure of animaging apparatus using an amorphous silicon thin film semiconductoraccording to one embodiment of the present invention, in which theapparatus uses photoelectric converters to take visible light andconvert that light into electrical signals.

In this X-ray imaging device, an operator can select between amoving-picture mode and a still-picture mode. It should be noted thatFIG. 1 does not show the fluorescent material used to convert the X-raysinto visible light, and although the present embodiment is describedwith reference to X-ray imaging, the present invention is not limited tosuch but should be understood as being equally applicable to instancesinvolving other types of radiation, such as, for example, alpha-rays,beta-rays or gamma-rays.

In FIG. 1, S1-1-S1-3 denote photoelectric converters arranged in atwo-dimensional array. T1-1-T1-3 denotes switches (that is, thin-filmtransistors, or TFT), G1-G3 denote gate wires that turn the TFT ON/OFFand M1-M3 denote signal lines. Each one of the photoelectric convertersis denoted by a photodiode and a capacitor connected in parallel, and isgiven a reverse bias. That is, the cathode terminal of the photodiode isgiven a positive (+) bias. The bias lines typically involve joint(common) wiring, which, for convenience, is not shown in FIG. 1. Thephotoelectrically converted electrical charges are then stored in thecapacitors. The entire assemblage of S1-1-S1-3, T1-1-T1-3, G1-G3, signallines M1-M3 and Vs lines is together called a radiation detectioncircuit or a photoelectric converter circuit. Reference numeral 102denotes a shift resister that applies pulses to the gate lines G1-G3 andcontrols the driving of the photoelectric converters S1-1-S1-3.Reference numeral 107 denotes a read-out circuit that amplifies andserially outputs the parallel signal output of the signal lines M1-M3inside the photoelectric converter circuit 101. Reference numeral 108denotes a signal processing circuit that processes the electricalsignals output from the read-out circuit 107.

Each of the photoelectric converters has a wavelength converter forconverting the radiation to visible light and a photoelectric converterproper that converts the visible light received into electrical signals.The matrix material of the wavelength converter is a material selectedfrom the group consisting of Gd₂O₃, Gd₂O₂S and CsI. The photoelectricconverter proper is composed mainly of amorphous silicon. Each of thephotoelectric converters is made of a material selected from the groupconsisting of amorphous silicon, gallium arsenic, mercury iodide andlead iodide. It should be noted that although for convenience ofexplanation the foregoing description uses the example of a 3×3-pixelphotoelectric converter, in reality many more photoelectric convertersare arrayed.

FIG. 2 is a circuit diagram showing the internal structure of theread-out circuit of FIG. 1. In FIG. 2, RES1-RES3 are reset switches thatreset the signal lines M1-M3, A1-A3 are amps that amplify the signals inM1-M3, CL1-CL3 are sample-hold capacitors that temporarily store thesignals amplified by amps A1-A3, Sn1-Sn3 are sample-hold switches, B1-B3are buffer amps, Sr1-Sr3 are switches for the serial conversion ofparallel signals, reference numeral 103 denotes a shift resister forapplying pulses to the switches Sr1-Sr3 for serial conversion, andreference numeral 104 denotes a buffer amp for outputting the seriallyconverted signals.

FIG. 3 is a schematic diagram showing the internal structure of thesignal processing circuit of the X-ray imaging apparatus of FIG. 1. Thesignal processing circuit 108 is composed of an AD converter thatconverts the analog output from the read-out circuit 107 to digitaloutput, FPN memories 1-4 for storing N=4 times' worth of dark output(FPN output), a light output memory 5 for storing light output byread-out after X-ray exposure, a CPU 6 that controls the signalprocessing circuit 108 and signal-processes the data stored in the FPNmemories 1-4 and the light output memory 5 in the ways described indetail below, and a shift resister 7 controlled by the CPU 6.

FIG. 4 is a timing chart of the operation of an imaging apparatusaccording to one embodiment of the present invention. In this example,there are 24 frames and the dark output is read four times as theprocess proceeds from left to right. The X-ray exposure interval shownin the timing chart of FIG. 7 is denoted by the reference symbol X inFIG. 4, and the read-out interval of FIG. 7 is shown as “H” in FIG. 4.“H” is the light output read-out interval. In FIG. 4, “X” and “H” arewritten at the 24th frame. “W” denotes the same time interval as theX-ray exposure period, but without the X-ray exposure; it is a “wait”interval in which the photoelectric converters “store darkness”, so tospeak. “K” and “F” denote a read-out after the wait interval, with thephotoelectric converters operating exactly as they do during interval“H”. “K” and “H” denote exactly the same read-out operation, the onlydifference being whether or not they are used as information foracquiring an X-ray image. “K” is called an “empty read” interval orempty read operation and “F” is called a FPN read interval or FPN readoperation. As information of acquiring an X-ray operation, “F” is used.

A description is now given of the operation of the X-ray imagingapparatus described above with reference to FIG. 4, in-which the darkoutput read operation is performed four times.

First, the X-ray imaging apparatus is caused to perform “W” and “K”operations alternately. At this time, at the actual imaging site, theradiologist (or X-ray technician, as the case may be) positions thatpart of the patient's body (that is, the subject) to be imaged atop thelight-sensing surface of the X-ray imaging apparatus. Since it is notknown when an X-ray will next be taken, the dark output information ofthe empty read “K” of this interval is stored in the signal processingcircuit 108. Once the radiologist or X-ray technician has properlypositioned the patient and completed preparation for imaging, the orderis given for the X-ray imaging apparatus to expose, at which point theX-ray imaging apparatus performs the wait operation “W” and the emptyread operation “K” (actually “F” but for convenience called “K”), thenshifts to the X-ray exposure interval “X”, and finally performs thelight output read-out operation “H”.

When imaging is finished, the four empty read operations “K” performedjust prior to the present read-out operation are used as FPN correctioninformation, denoted by “F” in FIG. 4. In this case, the empty readoperations of frame nos. 21, 22 and 23 are “F”. However, at the timeoperation commences, it cannot be determined in advance which empty readoperation “K” will be used for the FPN read operation “F”, so in essencethe “F” immediately preceding the “H” is used when the read operation“H” is finished by the exposure order issued by the radiologist or X-raytechnician.

Therefore, the information from “K” from the start of imaging must bestored in memory. However, since it is not known whether the exposurecommand will be issued immediately after operation begins, it makes nosense to store all the FPN in memory because doing so takes up a largeamount of memory capacity.

In order for the signal processing circuit 108 shown in FIG. 3 to storethe analog output from the X-ray imaging apparatus, the analog outputfrom the read-out circuit 107 must first must be converted into digitaloutput at the AD converter 11 and stored in the FPN memories 1-4 and thelight output memory 5. The AD converter output is connected in parallelto the data lines of each of the five memories described above. The CPU6 data bus is also connected to the data liens of each of the fivememories described above. At the same time, the CPU 6 address bus isconnected to the address lines of each of the five memories describedabove. The chronologically output N=4 times' worth of dark output (FPNoutput) is then stored in the respective memories by the output signals(G1-G4) output by the shift resister 7 which is controlled by the CPU 6.Each time the number of frames increases, the CPU 6 causes the shiftresister 7 shifts the output signals in the order G1→G2→G3→G4, andreturning once again to G1 output signal after the G4 output signal hasbeen output. Therefore, a continuous 4-time empty read operation “K”output is always being stored in the respective memories. In otherwords, even when all four memories have been written to, so long asthere is no exposure command issued new dark output data is successivelystored in the memories. After an exposure command has been issued andthe process proceeds to the X-ray exposure interval, the operation ofwriting data to the FPN memories is completed. The data that is storedin the FPN memories 1-4 then becomes the N=4 times' worth of dark outputdata immediately preceding the light output read-out operation “H”.

After the X-ray exposure interval is completed, the light output dataoutput by the read-out operation “H” is then stored in the light outputmemory 5 by a light output control signal from the CPU 6.

FIG. 5 is a schematic diagram showing steps in a process of averagingtwo continuous pieces of FPN data according to one embodiment of thepresent invention. FPN data corresponding to the pixels arranged in a2-dimensional array is stored in the respective FPN memories 1-4. TheCPU 6 then acquires FPN data for each of the respective pixels via thedata bus described above. The CPU 6 can easily average the four darkoutput operation values acquired as described above. The averagingprocess or operation referred to here means simply adding four times'worth of FPN to each of the respectively pixels and dividing the sum by4. The (average) value thus computed by the CPU may either be storedeasily in memory or written over any data already present in any of theFPN memories 1-4.

The averaged correction FPN data calculated by the CPU 6 as describedabove is then subtracted from each of the pixels. The light output dataand the correction FPN data are both acquired by the CPU 6 via the databus, with the latter then be subtracted fro the former to obtain acorrected result.

By the above-described process, an FPN-corrected, randomnoise-suppressed X-ray image data can be produced.

As described above, according to the imaging apparatus of the presentembodiment, when performing see-through imaging of a subject, theapparatus corrects the fixed pattern noise generated during the processof forming that see-through image of the subject and in so doing alsosuppresses increases in random noise generated by the photoelectricconverter circuit 101 and the read-out circuit 107, thereby providing aradiological imaging method and apparatus with improved reliability.

(Variation)

A description is now given of a variation of the embodiment describedabove, with reference to FIG. 6.

FIG. 6 is a schematic diagram showing steps in a process of averagingtwo continuous pieces of FPN data according to a variation of anembodiment of the present invention. FIG. 6 shows an example in whichthe dark output read-out operation is carried out twice. First, theX-ray imaging apparatus causes the “W” and “K” operations to beperformed, at which time, at the actual imaging site, the radiologist(or X-ray technician, as the case may be) positions that part of thepatient's body (that is, the subject) to be imaged atop thelight-sensing surface of the X-ray imaging apparatus. Unlike the imagingapparatus and method according to the first embodiment, in thisvariation thereof the dark output information of the empty readoperation “K” of this interval is not stored in memory. Once theradiologist or X-ray technician has properly positioned the patient andcompleted preparation for imaging, the order is given for the X-rayimaging apparatus to expose, at which point the X-ray imaging apparatusperforms the wait operation “W” and the empty read operation “K”(actually “F” but for convenience called “K”), then proceeds to the.X-ray exposure interval “X” and then proceeds to the light outputread-out operation “H”. Immediately after the light output read-outoperation is finished, the wait operation “W” and the read-out operation“H” are each performed twice, thus completing the imaging process. Thetwo empty read operations performed immediately after the light outputread-out operation “H” are stored in memory and, since they are used asdark output correction data (that is, FPN correction data), are denotedin FIG. 6 by the reference symbol “F”.

In the radiological imaging apparatus of the present embodiment, if thestandard deviation of the random noise included in the dark output imageobtained by the “dark output read-out operation” is σ1, then thestandard deviation included in the dark output image after carrying outsuch “dark output read-out operation” N times and averaging the resultsis σ1{square root}N.

If the standard deviation of the random noise included in the lightimage output obtained by the “light output read-out operation” is σ2,then the standard deviation σt of the random noise included in theultimate radiological image after calculating from the light imageoutput the dark output image using the above-described averaging isσt={σ2²+(σ1/{square root}N)²}^(1/2).

Here, the shot noise of the radiation included in the actual light imageoutput is not needed here and is ignored. The σ2 of the random noiseincluded in the light image output is the same as the σ1 included in thedark image output, so, if this is set simply at σ, then:σt=σ(1+1/N)^(1/2).

In the conventional case where N=1, random noise of σt=1.41 is improvedwith each averaging N such that, at N=2, σ=1.22, at N=3, σt=1.15, atN=4, σt=1.11.

Also, in the present invention, the dark image output read-out operationis carried out just before or just after the light image output, and sois not affected by the factors that cause fluctuations in fixed patternnoise such as the passage of time or changes in the ambienttemperature,.making it possible to provide a radiological imagingapparatus with very high correction accuracy.

The N=2 times' worth of FPN data “F” are afterwards averaged for each oftheir respective corresponding pixels to produce one frame of darkoutput data (FPN correction data). The averaged FPN correction data, issubtracted from the light output data of the light output read-outoperation “H” for each respective corresponding pixel to obtain anFPN-corrected, improved X-ray image data.

FIG. 8 shows experimental and theoretical measurements of noise levelversus number of averaging processes using the X-ray imaging apparatus.The radiological imaging apparatus is provided with a radiationdetection means in which 7.2 million pixels are arranged in a2-dimensional array, drive means, signal read-out means and signalprocessing means. In a first read-out operation, although in actualusage the apparatus reads the X-rays that pass through the subject, inorder to show the noise reduction effect here, read-out is performed ina completely dark environment. Similarly, a second read-out operation isalso conducted under the same blacked-out conditions, for a total ofeight measurements in all. In FIG. 8, the horizontal X axis representsthe number of averagings performed in the second read-out operation andthe vertical axis represents the standard deviation of 7.2 millionpixels after subtracting the second read-out operation data (averageddata) from the first read-out data. Incidentally, a conventional exampleis shown where the number of averagings is 1 on the horizontal axis. Inthis case, no averaging is performed because there is only one data.From FIG. 8, it is experimentally confirmed that making the number ofaveragings ≧2 reduces noise, ultimately converging at 1/{square root}2of noise level (3.8) when n=1. The extent of the improvement due toaveraging is greatest when the number of averagings is 2, producing a22.4 percent decrease in noise compared to a conventional example inwhich n=1 ({square root}2=41.4 percent increase). Where n=3, theincrease is 15.5 percent. As the number of averagings increases there isa convergence toward the theoretical value (3.8÷{square root}2=2.687),but increasing the number of averagings also increases the processingtime, so there is a trade-off between noise and speed. As can beunderstood from FIG. 8, as the number of averagings increases the noiselevel also decreases, but the effect lessens as the number of averagingsincreases. In an image including random noise, if there is a 10 percentincrease with respect to the convergence value (that is, theoreticalvalue) for that random noise, there is no major deterioration in picturequality, and moreover, is within the factory production S/N tolerancerange. In other words, with a maximum of five averagings (n=5), theobject of the present invention can be achieved, and conversely,performing greater than five averagings actually contributes todecreased utility. Accordingly, 2-5 averagings is an optimal range, inwhich noise is greatly reduced in little processing time.

As described above, according to the present variation, when performingsee-through imaging of a subject, the apparatus corrects the fixedpattern noise generated during the process of forming that see-throughimage of the subject and in so doing also suppresses any increase inrandom noise generated by the photoelectric converter circuit 101 andthe read-out circuit 107, thereby providing a radiological imagingmethod and apparatus with improved reliability.

According to the above-described radiological imaging apparatus andmethod, the dark output including FPN is subtracted from the lightoutput which includes the X-ray imaged so as to provide a radiologicalimaging method and apparatus with improved reliability.

Moreover, according to the radiological imaging apparatus and method ofthe present invention, the dark output containing FPN is acquired eitherjust before or just after imaging of the light output containing theX-ray image, so FPN correction accuracy is high.

Further, since FPN data, which has been averaged over multiple darkoutputs, is used to carry out FPN correction, increases in random noiseattendant upon FPN correction can also be suppressed, thus improving thequality of the radiological image.

As many apparently widely different embodiments of the present inventioncan be made without departing from the spirit and scope thereof, it isto be understood that the invention is not limited to the specificpreferred embodiments described above thereof except as defined in theclaims.

1-9. (canceled)
 10. A radiological imaging system for formingradiographic images, comprising: a radiation source; and a radiologicalimaging device including, a radiation detection unit having a pluralityof photoelectric converters arrayed in a matrix-like formation for thepurpose of detecting and converting radiation originated from saidradiation source into electrical signals, a read-out unit operativelyconnected to said radiation detection unit for reading the electricalsignals output by said radiation detection unit, and a signal processingunit operatively connected to said read-out unit for processing theelectrical signals read out by the read-out unit, wherein said signalprocessing unit performs a first read-out operation, in which theradiation detection unit is operated with applying the radiationoriginated from said radiation source, and a second read-out operation,in which the radiation detection unit is operated without applying theradiation selectively, and performs said second read-out operation atleast twice, averaging the results of said second read-out operation soas to obtain an average signal value thereof, and subtracting saidaverage signal value from a signal value obtained by said first read-outoperation so as to obtain a processed image.
 11. The radiologicalimaging system of claim 10, wherein said signal processing unit performssaid second read-out operation at least twice before performing saidfirst read-out operation.
 12. The radiological imaging system of claim10, wherein said signal processing unit performs said second read-outoperation at least twice after performing said first read-out operation.13. The radiological imaging system of claim 10, wherein each of theplurality of photoelectric converters comprises a wavelength conversionelement configured to convert said radiation into visible light and aphotoelectric conversion element configured to detect and convert saidvisible light into one or more electrical signals.
 14. A radiologicalimaging method implemented in a radiological imaging system including aradiation source, and a radiation detection unit having a plurality ofphotoelectric converters arrayed in a matrix-like formation fordetecting and converting radiation originated from the radiation sourceinto electrical signals, the method comprising: the first reading stepof reading the electrical signals output by the radiation detectionunit, in which the radiation detection unit is operated with applyingthe radiation originated from the radiation source; the second readingstep of reading the electrical signals output by the radiation detectionunit, in which the radiation detection unit is operated without applyingthe radiation, wherein said second reading step is performed at leasttwice; the averaging step of averaging the results of said secondreading step so as to obtain an average signal value thereof; and thesubtraction step of subtracting the average signal value from a signalvalue obtained at said first reading step.
 15. The radiological imagingmethod of claim 14, wherein said second reading step is performed atleast twice before said first reading step.
 16. The radiological imagingmethod of claim 14, wherein said second reading step is performed atleast twice after said first reading step.
 17. A program for controllinga radiological imaging system including a radiation source, and aradiation detection unit having a plurality of photoelectric convertersarrayed in a matrix-like formation for detecting and convertingradiation originated from the radiation source into electrical signals,the program comprising: code of the first reading step of reading theelectrical signals output by the radiation detection unit, in which theradiation detection unit is operated with applying the radiationoriginated from the radiation source; code of the second reading step ofreading the electrical signals output by the radiation detection unit,in which the radiation detection unit is operated without applying theradiation, wherein said second reading step is performed at least twice;code of the averaging step of averaging the results of said secondreading step so as to obtain an average signal value thereof; and codeof the subtraction step of subtracting the average signal value from asignal value obtained at said first reading step.
 18. The program ofclaim 17, wherein said second reading step is performed at least twicebefore said first reading step.
 19. The program of claim 17, whereinsaid second reading step is performed at least twice after said firstreading step.